Methods and systems for in vivo full-field interference microscopy imaging

ABSTRACT

According to one aspect, the invention relates to a system ( 101 ) for in vivo, full-field interference microscopy imaging of a scattering three-dimensional sample. It comprises a full-field OCT imaging system ( 130 ) for providing en face images of the sample, wherein said full-field OCT system comprises an interference device ( 145 ) with an object arm ( 147 ) intended to receive the sample and a reference arm ( 146 ) comprising an optical lens ( 134 ) and a first reflection surface ( 133 ), and an acquisition device ( 138 ) configured to acquire a temporal succession of two-dimensional interferometric signals (I 1 , I 2 ) resulting from interferences produced at each point of an imaging field; an OCT imaging system ( 110 ) for providing at the same times of acquisition of said two-dimensional interferometric signals, cross-sectional images of both the sample and a first reflection surface ( 133 ) of said full-field OCT imaging system ( 130 ); a processing unit ( 160 ) configured to determine a plurality of en face images (X-Y) of a plurality of slices of the sample, each en face image being determined from at least two two-dimensional interferometric signals (I 1 , I 2 ) having a given phase shift; determine from the cross-sectional images provided by the OCT imaging system ( 110 ) at the times of acquisition of each of said two two-dimensional interferometric signals (I 1 , I 2 ) a depth (z) for each en face image (X-Y) of said plurality of slices; determine a 3D image of the sample from said plurality of en face images of said plurality of slices of the sample and depths.

TECHNICAL FIELD

The present description relates to in vivo full-field interferencemicroscopy imaging methods and systems. It is applicable to in vivoimaging of randomly movable objects, and more particularly to in vivoimaging of ophthalmic tissues.

STATE OF THE ART

During its 25 years of development, optical coherence tomography (OCT)has become a powerful imaging modality (See for example “OpticalCoherence Tomography—Technology and Applications”—Wolfgang Drexler—JamesG. Fujimoto—Editors—Springer 2015). OCT is an interferometric technique,which can be seen as an “optical analogy” of ultrasound imaging. OCT hasapplications in a broad spectrum of areas, and in particular inapplications of the biomedical fields in ophthalmology, dermatology,cardiovascular field, gastroenterology.

In vivo tissues are involuntarily moving, and these movements have beenposing challenges for all OCT techniques throughout the history. Moreprecisely, movements lead to the appearance of misaligning, shifting anddoubling artifacts in the conventional scanning OCT images. Types of theartifacts are connected with the method of OCT, according to which allthe image pixels are not acquired at the same time, but rather byscanning point-by-point over the sample.

Desire to avoid these motion artifacts in the images motivated theprogress in OCT technology to achieve higher imaging speeds, whichresulted in spectral domain OCT (SD-OCT) (See for example L. An et al.“High speed spectral domain optical coherence tomography for retinalimaging at 500,000 A-lines per second”—Biomedical Optics Express 2, 2770(2011)) and more recently swept source OCT (SS-OCT) (See for example B.Potsaid et al. “Ultrahigh speed 1050 nm swept source/Fourier domain OCTretinal and anterior segment imaging at 100,000 to 400,000 axial scansper second”, Optics Express 18, 20029 (2010)), capable of imaging fasterthan 300,000 A-scans/s (1D profile). However, even at that speed ofscanning, OCT images are not immune to in vivo motion artifacts.

With the same goal to get images without motion artifacts, severalpublications and patents suggested software and hardware-based motioncompensation schemes (See for example M. Kraus et al. “Motion correctionin optical coherence tomography volumes on a per A-scan basis usingorthogonal scan patterns”, Biomedical Optics Express 3, 1182 (2012)).However, hardware-based solutions are bringing additional complexity tothe devices and are frequently bulky and expensive, while software basedare sample- and motion-specific, meaning that they can compensate only afew types of movements only of the particular objects.

A special case of OCT, called full-field OCT (FFOCT), uses a camera toacquire all the image pixels simultaneously without point-by-point orline-by-line scanning, and is, therefore, immune to the above-mentionedartefacts. The full-field OCT imaging technique is for example describedin the article “Full-field optical coherence tomography” by F. Harms etal. taken from the work “Optical Coherence Tomography—Technology andApplications”—pages 791-812—Wolfgang Drexler—James G.Fujimoto—Editors—Springer 2015. The full-field OCT imaging technique isalso described in the French patent application FR2817030.

The full-field OCT imaging technique is based on the use of the lightbackscattered by a sample when it is illuminated by a light source withlow coherence length, and in particular the use of the lightbackscattered by the microscopic cell and tissue structures in the caseof a biological sample. This technique exploits the low coherence of thelight source to isolate the light backscattered by a virtual slice depthwise in the sample. The use of an interferometer makes it possible togenerate, by an interference phenomenon, an interference signalrepresentative of the light originating selectively from a given sliceof the sample, and to eliminate the light originating from the rest ofthe sample. More specifically, in order to obtain a single 2D FFOCTimage, acquisition of several (typically 2 to 5) direct images on thecamera is performed. Each of these direct images is acquired with aparticular interference phase, which is set by a precisely positionedmirror with a piezo element (PZT) in the reference arm of theinterferometer. Post-processing of these direct images with theparticular phases allows to retrieve an FFOCT image.

Besides the above-mentioned immunity to the scanning artifacts, FFOCTprovides higher lateral resolution by using high numerical apertureobjectives than OCT since typical OCT uses relatively low NA objectivesdue to the requirements of large depth of field. It gives similar axialresolution by using the cheap broadband spatially incoherentillumination sources.

However, current 2D imaging scheme of FFOCT is practical for the staticsamples (or for in vivo samples in the moments of no or low movements),as any motion of the sample may shift the pre-determined phases anddegrades the FFOCT signal or even destroys the FFOCT image. The schemesfor 3D imaging are not applicable either for in vivo imaging, aslocations (X, Y, Z) of the captured 2D images are becoming unknown,therefore making construction of the 3D image impossible. As a result,up to now applications of FFOCT were almost entirely limited to thestatic ex vivo samples.

The present description is related to devices and methods which have theadvantages of full-field optical coherence tomography, and which at thesame time can perform imaging of constantly moving in vivo objects.

SUMMARY

According to a first aspect, the present description relates to a methodfor in vivo, full-field interference microscopy imaging of a scatteringthree-dimensional sample comprising:

-   -   disposing the sample in an object arm of an interference device        of a full-field OCT imaging system, wherein said interference        device further comprises a reference arm with an optical lens        and a first reflection surface;    -   producing, at each point of an imaging field, an interference        between a reference wave obtained by reflection of incident        light waves on an elementary surface of the first reflection        surface corresponding to said point of the imaging field and an        object wave obtained by backscattering of incident light waves        by a voxel of a slice of the sample at a given depth, said voxel        corresponding to said point of the imaging field,    -   acquiring, using an acquisition device of said full-field OCT        imaging system, a temporal succession of two-dimensional        interferometric signals resulting from the interferences        produced at each point of the imaging field;    -   storing, for each two-dimensional interferometric signal, a time        of acquisition;        -   providing, at each time of acquisition of the            two-dimensional interferometric signals, cross-sectional            images of both the sample and said first reflection surface            of said full-field OCT imaging system using an OCT imaging            system;    -   determining a plurality of en face images of a plurality of        slices of the sample, each en face image being determined from        at least two two-dimensional interferometric signals having a        given phase shift;    -   determining from the cross-sectional images provided by the OCT        imaging system at the times of acquisition of each of said two        two-dimensional interferometric signals a depth for each en face        image of said plurality of slices;    -   determining a 3D image of the sample from said plurality of en        face images of said plurality of slices of the sample and        depths.

In the present specification, “en face images” are images determined ina plane (“X-Y” plane) perpendicular to an optical axis of the object arm(also referred to as sample arm). “En face images” are also referred as“X-Y images” or “FFOCT signal” in the present specification.

“Cross-sectional images” are images (1D or 2D) determined in a planethat contains an optical axis of the object arm. Cross-sectional imagesare also referred to as “X-Z images” in the present specification;however, they are not limited to a particular plane and may bedetermined in any plane perpendicular to the “X-Y” plane.

An “optical lens” in the present specification refers to any opticaldevice that focuses or disperses light by means of light refraction. An“optical lens” thus encompasses both conventional optical lenses(convex, plano-convex, doublets, etc.) as other imaging systems (e.g.microscope objectives).

The imaging method thus described makes it possible to preciselydetermine the depth of the slice that is imaged by the FFOCT imagingsystem, even when imaging in vivo samples having natural movements. Thisis made possible by providing simultaneous acquisition oftwo-dimensional interferometric images using the FFOCT imaging systemand the cross-sectional images provided by the OCT imaging system.

In vivo natural movements of the object can thus be used for 3D imaging,meaning that we take advantage of an effect that most of the methods tryto eliminate or to overcome.

According to one or a plurality of embodiments, determining a depth foreach en face image of said plurality of slices of the sample comprisesdetermining a relative axial position of said first reflection surfaceand at least one identified structure of the sample in thecross-sectional images provided by the OCT imaging system.

Practically speaking, the plurality of en face images of the pluralityof slices of the sample may be determined within an explored volume ofthe sample. Depth of an en face image of a slice is determined from OCTimages from the difference between the axial location of the detectedreference mirror peak and the axial location of any sample peak. It isnot important, which peak of the sample is used, but typically, thebrightest peak may be used. However, the same sample peak will be usedthroughout one volume acquisition, so that relative depths of the enface slices are correct, and a 3D image can be determined.

According to one or a plurality of embodiments, said full-field OCTimaging system and said OCT imaging system being mounted on a movingplatform, the method further comprises moving said platform at leastalong the optical axis (Z) of the object arm to determine said pluralityof en face images.

According to one or a plurality of embodiments, the method furthercomprises moving said platform along at least one of the directions (X,Y) perpendicular to the optical axis of the object arm. It is thuspossible to stack cross-section images both axially and laterally andallow the formation of larger 3D volume (e.g. by image registration).

According to one or a plurality of embodiments, said object arm beingmounted on a moving platform, the method further comprises moving saidplatform along an optical axis of the object arm to determine saidplurality of en face images.

According to one or a plurality of embodiments, natural in vivomovements of the sample are used to determine said plurality of en faceimages. There is no need to move any of the platforms of said object armor said full-field OCT imaging system and said OCT imaging system.

According to one or a plurality of embodiments, e.g. for cornea imaging,the object arm further comprises an optical lens, e.g. a microscopeobjective. The depth of focus of such optical lens is much smaller thanthe depth of focus of the eye. As a result, when the relative positionof the sample arm and the sample is changed, the method furthercomprises moving the reference arm along an optical axis of saidreference arm to compensate for defocus, i.e. keep a coherence planewithin the depth of focus of the sample arm microscope objective. As amatter of fact, when moving from one medium to a second one, e.g. airand eye, a shift appears between the focus and the position thatequalizes the optical paths in both arms. This defocus needs to becompensated.

According to one or a plurality of embodiments, e.g. for retina imaging,the depth of focus is high and there is no need to compensate fordefocus when the relative position of the sample arm and the sample ischanged.

According to one or a plurality of embodiments, the method furthercomprises position shifting said first reflection surface of thereference arm of the full-field OCT imaging system to provide said phaseshift between said at least two two-dimensional interferometric signals.These embodiments suppose that the natural movements of the sample areslow during the time of acquisition of the at least two two-dimensionalinterferometric signals. Typical acquisition time is 1-10 ms.

According to one or a plurality of embodiments, the method furthercomprises selecting in said temporal succession of two-dimensionalinterferometric signals acquired by the acquisition device, said atleast two-dimensional interferometric signals having said phase shift,wherein the phase shift results from in vivo movements of the sample.

Here again, the natural movements of the in vivo sample are used for enface imaging, meaning that we take advantage of an effect that most ofthe methods try to eliminate or to overcome.

The different embodiments of the imaging method according to the firstaspect of the present description can be combined with one another.

According to a second aspect, the present description relates to asystem for in vivo, full-field interference microscopy imaging of ascattering three-dimensional sample, configured for implementing one ora plurality of embodiments of the method according to the first aspect.

According to one or a plurality of embodiments, the system according tothe second aspect comprises:

-   -   a full-field OCT imaging system for providing en face images of        the sample, wherein said full-field OCT system comprises:        -   an interference device comprising an object arm intended to            receive the sample and a reference arm comprising an optical            lens and a first reflection surface, wherein said object arm            and said reference arm are separated by a beam splitter and            wherein the interference device is adapted to produce, when            the sample is disposed on the object arm of the interference            device, at each point of an imaging field, an interference            between a reference wave obtained by reflection of incident            light waves on an elementary surface of the first reflection            surface corresponding to said point of the imaging field and            an object wave obtained by backscattering of incident light            waves by a voxel of a slice of the sample at a given depth,            said voxel corresponding to said point of the imaging field,        -   an acquisition device configured to acquire a temporal            succession of two-dimensional interferometric signals            resulting from the interferences produced at each point of            the imaging field, an OCT imaging system for providing at            the same times of acquisition of said two-dimensional            interferometric signals, cross-sectional images of both the            sample and said first reflection surface of said full-field            OCT imaging system;    -   a processing unit configured to:        -   determine a plurality of en face images of a plurality of            slices of the sample, each en face image being determined            from at least two two-dimensional interferometric signals            having a given phase shift;        -   determine from the cross-sectional images provided by the            OCT imaging system at the times of acquisition of each of            said two two-dimensional interferometric signals a depth for            each en face image of said plurality of slices;        -   determine a 3D image of the sample from said plurality of en            face images of said plurality of slices of the sample and            depths.

The advantages stated for the imaging method can be transposed to theimaging system according to the second aspect of the presentdescription.

According to one or a plurality of embodiments, said first reflectionsurface of the reference arm of the full-field OCT imaging system isposition shifted to provide said optical path difference between said atleast two-dimensional interferometric signals.

According to one or a plurality of embodiments, said first reflectionsurface of the reference arm of the full-field OCT imaging system isfixed and the processing unit is further configured to select in saidtemporal succession of two-dimensional interferometric signals acquiredby the acquisition device, said at least two-dimensional interferometricsignals having said given optical path difference, wherein the opticalpath difference results from in vivo movements of the sample.

According to one or a plurality of embodiments, said object arm of thefull-field OCT imaging system further comprises an optical lens.

According to one or a plurality of embodiments, said optical lens of thereference arm and/or object arm is a microscope objective.

According to one or a plurality of embodiments, said reference armand/or object arm of the full-field OCT imaging system can be moved withrespect to said beam splitter of the interference device of saidfull-field OCT imaging system (along each optical axis of said referencearm and object arm).

According to one or a plurality of embodiments, the system furthercomprises a moving platform, wherein said full-field OCT imaging systemand said OCT imaging system are mounted on said moving platform.

According to one or a plurality of embodiments, the OCT imaging systemis a spectral domain OCT imaging system or a swept-source OCT imagingsystem, or a time-domain OCT imaging system.

The different embodiments of the imaging system according to the presentdescription can be combined with one another.

Different features and embodiments of the various aspects of the presentdescription can also be combined with one another.

BRIEF DESCRIPTION OF THE FIGURES

Other advantages and features of the imaging technique presentedhereinabove will become apparent on reading the following detaileddescription, with reference to the figures in which:

FIGS. 1A and 1B are schemes of systems according to embodiments of thepresent description;

FIGS. 1C and 1D show exemplary light sources spectra of the OCT sourceand the FFOCT source and blocking parts of such spectra with filters ofthe system, according to embodiments of the present description;

FIGS. 2A, 2B are flow diagrams of embodiments of an imaging methodaccording to the present description and images to illustrate steps ofsome of these embodiments;

FIGS. 3A, 3B are flow diagrams of further embodiments of an imagingmethod according to the present description and images to illustratesteps of some of these embodiments;

FIG. 4A is an example of a cross-sectional image of the reference mirror(no sample visible) obtained using an exemplary OCT imaging system of animaging system according to the present description and FIG. 4Brepresents the curve of variation of intensity along the vertical lineof FIG. 4A;

FIG. 5A is an example of a cross-sectional image of the reference mirror(with cornea sample visible) obtained using an exemplary OCT imagingsystem of an imaging system according to the present description andFIG. 5B represents the curve of variation of intensity along thevertical line of FIG. 5A;

FIG. 6, is a graph illustrating required phase-shifting by the in vivomovements;

FIG. 7 shows FFOCT images of in-depth layers of the in vivo human corneaacquired using phase shifting by the natural eye movements (no move ofthe reference mirror).

FIG. 8 shows FFOCT images of in-depth layers of the in vivo human cornea(stroma) acquired using phase shifting by the natural eye movements (nomove of the reference mirror) at different camera exposure times.

DETAILED DESCRIPTION

Systems

FIGS. 1A and 1B show respectively two embodiments 101, 102 of a systemfor in vivo, full-field interference microscopy imaging according to thepresent description. The system 101 is suitable for implementing amethod for 3D imaging of in vivo moving samples, and particularly, butnot limited to, the anterior part 11 (cornea) of the in vivo eye. Thesystem 102 is suitable for implementing a method for 3D imaging of invivo moving samples, and particularly, but not limited to, the posteriorpart 13 (retina) of the in vivo eye.

The system 101 shown in FIG. 1A comprises two imaging systems, afull-field OCT (“FFOCT”) imaging system 130 and an optical coherencetomography (“OCT”) imaging system 110 and at least one processing unit160. The FFOCT imaging system enables to get “en face” images of amoving in vivo sample 11, i.e. images of in-depth sections of thesample, and the optical coherence tomography (“OCT”) imaging system 110provides information about the position of the sample in an axial (Z)direction, e.g. along an optical axis. The system 101 may furthercomprise a moving platform 150, e.g. with one or several motors, onwhich the FFOCT imaging system 130 and the OCT imaging system 110 aremounted. The moving platform 150 is capable of translating jointly theFFOCT imaging system and the OCT imaging system in all X, Y and Zperpendicular directions.

The FFOCT imaging system 130 of FIG. 1A comprises an interference device145 and an acquisition device 138 connected to said at least oneprocessing unit 160.

According to one embodiment, the interference device 145 comprises abeam splitter element 135, for example a non-polarizing splitter cube,making it possible to form two arms, a reference arm 146 with opticalaxis Δ_(R), and an object arm 147 with an optical axis Δ₀. In FIG. 1A,the optical axis Δ₀ of the object arm defines the Z axis and the opticalaxis Δ_(R) of the reference arm defines the X axis. The reference arm146 comprises a reflection surface 133. The reflection surface 133 maybe flat; it is for example a metallic mirror, a neutral density (ND)filter glass, or simply a glass plate. The object arm 147 is intended toreceive, in operation, the three-dimensional scattering sample 11, avolume of which it is a desire to produce a tomographic image.

In the embodiment of FIG. 1A, the reflection surface 133 is mounted on apiezo electric stage (PZT) 132 for phase modulation; such phasemodulation may be used in one embodiment of the method according to thepresent description, and may not be used in another embodiment, as itwill be described further.

The interference device is adapted to produce optical interferencesbetween, on the one hand, reference waves obtained by reflection oflight emitted by a light source 141, spatially incoherent or of lowcoherence length, by each elementary surface of the reflection surface133 of the reference arm 146 and, on the other hand, of the object wavesobtained by backscattering of the light emitted by the same source byeach voxel of a slice of a sample 11 depth wise in the sample, thesample 11 being disposed on the object arm 147, said voxel and saidelementary surface corresponding to the same point of the imaging field.

The light source 141 is a source that is spatially incoherent and of lowtemporal coherence length (in practice, in a range from 1 to 20micrometers), for example a thermal light source (e.g. halogen lamp) ora LED. According to one or more exemplary embodiments, the light source141 can form part of the FFOCT imaging system 130, as in the example ofFIG. 1A, or can be an elemental external to the imaging system, theFFOCT imaging system 130 being configured to work with light wavesemitted by the source. An optical system 140 may be used to realize aKohler-like illumination. In operation, light emitted by the lightsource 141 is reflected by a dichroic mirror 139 and reachesbeam-splitter element 135 of the interference device 145.

The acquisition device 138 allows the acquisition of at least onetwo-dimensional interferometric signal resulting from the interferencesbetween the reference waves and the object waves.

The acquisition device 138 is for example an image sensor, of CCD(Charge-Coupled Device) or CMOS (Complementaritymetal-oxide-semiconductor) camera type. This acquisition device iscapable of acquiring images at a high rate, for example with a frequencycomprised between 100 Hz and 1000 Hz, or higher. Depending on thedynamics of the sample studied, and more specifically the dynamics ofthe movements within the sample, it is possible to use the camerasoperating from a few Hz up to several KHz.

The processing unit 160 is configured to execute at least one step ofprocessing of at least one two-dimensional interferometric signalacquired by the acquisition device 138 and/or at least one step of imagegeneration in accordance with at least one of the imaging methodsaccording to the present description, in order to generate at least oneimage of the sample slice.

In one embodiment, the processing unit 160 is a computing devicecomprising a first memory CM1 (not represented) for the storage ofdigital images, a second memory CM2 (not represented) for the storage ofprogram instructions and a data processor, capable of executing programinstructions stored in this second memory CM2, in particular to controlthe execution of at least one step of processing of at least onetwo-dimensional interferometric signal acquired by the acquisitiondevice 138 and/or of at least one step of image computation inaccordance with at least one of the imaging methods according to thepresent description.

The processing unit can also be produced in integrated circuit form,comprising electronic components suitable for implementing the functionor functions described in this document for the processing unit. Theprocessing unit 160 can also be implemented by one or more physicallydistinct devices.

In the example of FIG. 1A, the interference device is a Linnikinterferometer and comprises two optical lenses 134, 142, e.g.microscope objectives, arranged on each of the reference and objectarms. The microscope objectives 134, 142 may have a relatively highnumerical aperture (typically ˜0.3), which at the same time providesrelatively large field-of-view (typically ˜1 mm). The reflection surface133 is thus located at the focus of the objective 134 of the referencearm and the sample 11 is intended to be positioned at the focus of theobjective 142 of the object arm; more specifically, the layer ofinterest of the sample is intended to be positioned at the focus of theobjective 142. Other types of interferometers can be envisaged for theimplementation of the methods according to the present description, andin particular but without limitation Michelson interferometers.

In the example of FIG. 1A, the microscope objective 142 of the objectarm 147 is mounted on a motorized platform 143, movable along thedirection of an optical axis of said object arm (Z axis), i.e, closer orfurther from the sample 11. Both the reflective surface 133 andmicroscope objective 134 of the reference arm 146 are mounted on anothermotorized platform 131, which can move along the direction of an opticalaxis of said reference arm (X axis).

At the output of the interferometer 145, there may be an opticalspectral filter 136 and an optic lens 137, for example an achromaticdoublet, whose focal length is adapted to allow a suitable sampling ofthe sample 11 by the acquisition device 138, and which makes it possibleto conjugate the planes situated at the foci of the two objectives and adetecting surface of the acquisition device 138. The acquisition device138 thus acquires the interference signals produced by the interferencedevice. In order to not limit the resolution permitted by the microscopeobjectives 134 and 142, the choice of the focal length of the optic 137will be in line with the Shannon sampling criterion. The focal length ofthe optic 137 is for example a few hundreds of millimeters, typically300 mm.

The optical spectral filter 136 advantageously transmits the wavelengthsof the light source 141, while blocks the wavelengths of the OCT source112, as further described below.

Glass plates or, so called dispersion compensation blocks (notrepresented in FIG. 1A), may be provided on each of the arms tocompensate the dispersion.

The OCT imaging system 110 comprises a spatially coherent light source112, a detector 113 and an interference device with a beam splitterelement 114 that defines a reference arm and an object arm of theinterference device of the OCT imaging system. Typically, the spatiallycoherent light source 112 can be a superluminescent diode (SLD), forexample in case of Spectral-Domain OCT or Time-Domain OCT, or a sweptlaser source. Typically, the detector 113 can be a device directlyconverting incident optical power into an electrical signal, for examplea photodiode, in case of Time-Domain-OCT or Swept-source OCT, or aspectrometer, in case of Spectral-Domain OCT.

The light from the source 112 is collimated into a fiber 118 and issplit by the beam splitter element 114 into two fibers 121 (object arm)and 120 (reference arm). In operation, after going through the fiber120, light passes through a lens 115, a dispersion compensation plate116, which can be rotated, and reaches a reflecting surface 117, forexample a metalized mirror. After going through the fiber 121, lightreaches a transverse scanning mechanism 111, which can scan the beam in2D (X-Y) directions. Then light beam passes though an optical filter122, passes though the dichroic mirror 139 and is split into the FFOCTreference arm 146 and the FFOCT sample arm 147 by the beam splitter 135.

Optical filter 122 is chosen in order to allow a light beam issued fromthe OCT source 112 to propagate in both the OCT reference arm, the FFOCTreference arm and the FFOCT sample arm but to block light from the FFOCTsource 141; on the other hand, optical filter 136 blocks the light beamissued from the OCT source 112 and pass the light from the FFOCT source.

Functionalities of the optical filters 122 and 136 are further describedin relation with FIGS. 1C, 1D. In the example shown in FIG. 1C, theoptical filter 122 may block (dashed lines) wavelengths of the OCT lightsource 112 that are below a given wavelength λ_(Filter122) which isabove the highest wavelength λ_(FFOCTmax) used in the FFOCT imagingsystem. On the other hand, as illustrated in FIG. 1D, optical filter 136in the FFOCT imaging system may block wavelengths of the OCT lightsource which are above a given wavelength λ_(Filter136) which is abovesaid highest wavelength λ_(FFOCTmax) used in the FFOCT imaging system.As a result, none of the OCT light reaches the acquisition device 138.

Obviously, FIGS. 1C and 1D only represent an example of functionalitiesof the optical filters 122, 136. Many other configurations are possibleas long as none of the OCT light reaches the acquisition device 138.

In a preliminary step, the optical pathlength of the OCT arm from thebeam splitter 114 to the mirror 117 (reference arm) may be matched withthe optical pathlength from the beam splitter 114 to the mirror 133 inthe FFOCT reference arm 146. Matching of the optical pathways of the OCTand

FFOCT reference arms may be achieved in a simple way. In real time, welook at the OCT images. If the mirror of the FFOCT reference arm is notvisible on the OCT images, then reference arms of OCT and FFOCT systemsare not matched. We extend the reference arm of the OCT imaging systemuntil the mirror of the FFOCT reference arm is visible on the OCTimages.

In operation, back-reflected light from the reflecting surface 133 inthe reference arm 146 of the FFOCT imaging system combines at the beamsplitter 135 with the back-reflected light from the different layers ofthe sample. Beam splitter 135 again divides the light into two parts:the reflected part is blocked by the filter 136 (as explained inrelation with FIG. 1D) and the transmission part passes though thedichroic mirror 139, filter 122, fiber 121. Then this light beam mixeswith the back-reflected light coming from the fiber 120 and, afterpassing through the fiber 119, is collected by the detector 113. Thedetector 113, for example a spectrometer, is configured to record the,so called, A-scan—a 1D profile, containing information about thereflectivity at different depths of the imaged object. Further, itcollects information about the position of the reference mirror 133 ofthe reference arm 146 of the FFOCT imaging system. By scanning the beamwith the scanning mechanism 111, it is possible to acquire 2D and 3Dreflectivity images.

The OCT imaging system may be a Spectral-Domain OCT (the detector 113 isa spectrometer) but it can be also a Time-Domain OCT or a Swept-SourceOCT.

The OCT imaging system may also provide information about the speed ofthe sample, based on several consecutive positions of the sample and thetime interval between them. Information about the instantaneous speed ofthe sample can be useful to predict its future movement (e.g. if thesample in the first moment is moving in a rapid way in Z direction, wecan expect that in the next moment it will continue to move in the samedirection).

As it will be further explained below, embodiments of the methodaccording to the present description use the above-mentioned OCT imagingsystem for obtaining information about the positions of the differentlayers of interest of the sample 11 and the position of the referencemirror 133 of the reference arm 146 of the FFOCT imaging system.

The system 102 shown in FIG. 1B is similar to that of FIG. 1A with minordifferences. In the embodiment of FIG. 1B, the role of the microscopeobjective in the sample arm 147 of the FFOCT imaging system is performedby the cornea 11 and the lens 12 of the eye. Additionally, an adaptivelens 148, for example a liquid lens, and a rotating glass plate 149 maybe inserted in the reference arm to compensate for the aberrations andthe dispersion mismatch introduced by the eye. Due to the typicallylarge depths of focus of the lens 12 of the eye, imaging of retinallayers of different depths can be performed without correcting fordefocus and, therefore, without moving the reference arm 146, contraryto the device shown in FIG. 1A. In all the other aspects, the system maybe similar to that of the embodiment shown in FIG. 1A.

3D Imaging Methods

FIG. 2A is a flow diagram of embodiments of an imaging method accordingto the present description. It can be implemented using a system asshown in FIG. 1A for example. The described methods are suitable for 3Dimaging of in vivo moving samples, and particularly, but not limited to,the anterior part 11 (cornea) of the in vivo eye.

FIG. 2B illustrates images acquired both by the OCT imaging system andthe FFOCT imaging system during the different steps shown in FIG. 2A.

Steps of FIGS. 2A, 2B illustrate acquisition of 3D images.

In step 201, images from the two devices, the OCT imaging system and theFFOCT imaging system, are obtained and displayed. FFOCT images can beobtained with either modulated PZT or static PZT, as it will bedescribed further. In corresponding step 201 in FIG. 2B, OCT image 221shows the mirror 133 of the reference arm of the FFOCT imaging system.On images 222 and 228 there is only the camera noise because defocuscorrection is not performed yet or/and optical pathways of the sampleand reference arms are not matched.

In step 203, it is checked whether the corneal layers are visible in theOCT images.

If NO, as shown in image 226 of FIG. 2B, the whole device 150 may bemoved along X, Y and Z axes (step 204). For example, it can be moved byan operator until the required layers are observed on the screen.

If YES, as shown in image 227 of FIG. 2B, the OCT image of the referencemirror 133 may be overlapped with a corneal layer (the reference corneallayer) by moving the whole device 150 along X and Z axes (step 205). Thereference corneal layer may be any layer, although, typically, a layerproviding a bright peak may be chosen. In FIG. 2B, image 229 shows theimage of the reference mirror that overlaps the image of a referencecorneal layer. At the same time, a fuzzy FFOCT image 230 can be seenbecause of lack of defocus correction.

A FFOCT image solely does not contain information about the location inthe sample, where the image was captured. OCT imaging system 110, usedin combination with FFOCT bridges this gap by providing X,Y,Zcoordinates of the captured image. Stack of 2D FFOCT images eachaccompanied with their locations can be grouped to form a 3D image. Moreprecisely, the method of 3D image acquisition 209 is described below.

In a first implementation (210), only the microscope objective 142 ismoved by the motor below the sample arm 147. At the same time thereference arm 146 is moved further from (or closer to) the beam splitter135 to compensate for the optical path mismatch between the sample arm147 and the reference arm 146.

In the second implementation (211), the whole device 150 is moved by themotor 101 closer to (or further from) the sample 11. At the same timethe reference arm 146 is moved further from (or closer to) the beamsplitter 135 to compensate for the optical path mismatch between thesample arm 147 and the reference arm 146.

In the third implementation (212), only the reference arm 146 is movedfurther from the beam splitter 135 to compensate for the opticalmismatch (defocus) between the sample arm 147 and the reference arm 146.Extent of the reference arm movement depends on the instantaneous sampleposition (or depth in the sample). Changes in the sample position (ordepth) are governed solely by in vivo sample movements.

In all implementations, individual en face images of slices at differentdepths in the sample are recorded according the methods described below.At the same time the position (X, Y, Z) of the slice corresponding toeach 2D image is recorded by the OCT imaging system 110. By having theposition information for each 2D image, 2D images can be repositioned inorder to form a 3D image.

Determination of the depth of each slice for which an en face image isacquired is made by storing (213) the times when those images areacquired. Acquisition is stopped when desired (214). In step 215, we usethe positional information from OCT images at the different times thathave been stored to realign 2D OCM images, i.e. images obtained by theFFOCT device and form a 3D corneal image.

Examples of 2D cross-section images and a 1D images (A-scans) used forposition detection are shown in FIG. 4A, 4B and FIG. 5A, 5B.

FIGS. 4A, 4B, 5A, 5B illustrate how the OCT imaging system is used todetermine the depth of a slice that is imaged by the FFOCT imagingdevice. It also shows the compensation for defocus.

FIGS. 4A and 4B show respectively the 2D cross-sectional image andcorresponding profile (A-scan) obtained using the OCT imaging systemwhen the sample is not introduced. FIGS. 5A and 5B show respectively the2D cross-sectional image and corresponding profile obtained using theOCT imaging system when the sample is introduced in the sample arm andis within the field of view of the OCT imaging device. The depth of aslice is measured relatively to the non-defocus corrected position of areference, corresponding to a black vertical line at 0 depth in FIGS.4A, 4B, 5A, 5B). For example, a reference layer is the top layer of thecornea.

On image 231 the very top layer of the cornea (cornea is shown inbracket) overlaps with the reference mirror (shown by an arrow). Thisposition corresponds to the “0” position in FIGS. 4A, 5A, i.e. when thedifference between the corneal top layer position and the non-defocuscorrected reference mirror position equals zero. As a result, nodefocusing correction is applied and we get the image of the cornealsurface 232.

On image 233, the corneal top layer is shifted up (on the image)relatively to the non-defocus corrected reference arm position. As aresult, a non-zero depth is measured. Based on this depth, reference armis shifted down (on the image) from the non-defocus corrected referenceposition. As a result, we get image 234 from the corneal layer, which inOCT image overlaps with the reference mirror image.

On image 235 everything is repeated as in the step before. Cornea isshifted up again and reference arm with mirror is shifted down again, asa result providing us with the FFOCT image from the deep cornea 236.

The embodiments described above are proposed to be used for imaging invivo moving samples and, particularly, the anterior part of the in vivoeye.

Embodiments of the method described below can also be used for imagingvarious in vivo samples, but the focus is, particularly, on theposterior part of the in vivo eye. FIG. 3A is a flow diagram ofembodiments of an imaging method according to the present description.It can be implemented using a system as shown in FIG. 1B for example.The described methods are suitable for imaging of in vivo movingsamples, and particularly, but not limited to, the posterior part 13(retina) of the in vivo eye.

FIG. 3B illustrates images acquired both by the OCT imaging system andthe FFOCT imaging system during the different steps shown in FIG. 3A.

In step 301, acquisition starts (Acquisition comprises the processing toobtain images) and images are displayed from the two devices, the OCTimaging system and the FFOCT imaging system. FFOCT acquisition can bedone with either modulated PZT or static PZT, as described below. Instep 303, it is checked whether the retinal layers are visible in theOCT images. On images 322 and 326 there is only the camera noise becausedefocus correction is not performed yet or/and optical pathways of thesample and reference arms are not matched.

If NO, as shown in image 324 of FIG. 3B, the whole device 150 is movedalong X, Y and Z axes (step 304). If YES, as shown in image 325 of FIG.3B, the OCT image of the reference mirror 133 is overlapped with theretinal layer of interest by moving the whole device 150 along X, Y andZ axes (step 305).

At that stage, optical path length is matched between mirror 133 and anyof the retinal layers, as illustrated in OCT images 327 or 329, FIG. 3Band corresponding FFOCT images 328, 330.

Then, 3D image acquisition 309 is started.

In a first implementation (310), only the reference arm 146 is moved bythe motor below.

In a second implementation (311), the whole device 150 is moved by themotor 101 closer to (or further from) the sample 11.

In a third implementation, none of the motors are moved and creation ofthe 3D stack is achieved by the in vivo movements of the sample.

As for FIGS. 2A, 2B, individual 2D cross-sectional images of the movingin vivo sample are recorded according to the mentioned above method.Device may be adjusted to get the maximum FFOCT signal on thefrequencies of the typical sample movements, if in vivo movements of thesample are used. At the same time the position (X, Y, Z) of the samplecorresponding to each 2D image is recorded by the OCT imaging system 110(steps 313, 314, 315 and corresponding images 331—336 on FIG. 3B). Byhaving the position information for each 2D image, 2D images can berepositioned in order to form a 3D image.

Determination of En Face Images

In order to extract an FFOCT image from the direct camera images aphase-shifting scheme is required.

In a first embodiment of the present description, a standard FFOCT imageretrieval method is used, according to which phase-shifting is providedby modulating the piezo element (PZT) 132. This embodiment is useful forthe case of slowly moving samples (their movement during the typicaltime of image acquisition should be <<π phase shift) or for thefast-moving samples in the moments of no movements. FFOCT image can beextracted from the 2, 4 or 5 direct images depending on the scheme.

For example, for 2 direct images:

${I\;\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack \left( {{\phi\left( {x,y} \right)} + \Psi} \right\rbrack \right\}}}} \right.}$

Where:

-   -   ϕ is the phase difference between the sample signal and the        reference signal;    -   ψ is the phase shift induced by PZT    -   I₀ is the photon flux of the illumination;    -   R_(ref)(x, y)≈const is the reflectivity of the reference, which        is spatially uniform;    -   R_(sum)(x, y) is the reflectivity of the sample structures        within the coherence volume, which is the plane of interest;    -   R_(inc)(x, y) is the reflectivity of all the other structures        that are out of the coherence volume and other stray        reflections.

Two phase-shifted images are:

${I_{1}\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)}{R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack {{\phi\left( {x,y} \right)} + 0} \right\rbrack}}} \right\}}$${I_{2}\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack {{\phi\left( {x,y} \right)} + \pi} \right\rbrack}}} \right\}}$

By subtracting the two images and taking the module we get the FFOCTimage or “FFOCT signal”.

${{{I_{1}\left( {x,y} \right)} - {I_{2}\left( {x,y} \right)}}} = {{I_{0} \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack {\phi\left( {x,y} \right)} \right\rbrack}}}$

Having a phase-shift between the two consecutive direct camera framesequals π (in a 2-phase-shifting scheme) enables to get the highestpossible FFOCT signal.

In a second embodiment of the present description, the image retrievalmethod used relies on the in vivo natural movements of the sample.

The applicants have shown that in ophthalmic tissue imagingapplications, for example, natural eye movements introduce phase changesbetween consecutive direct images, which can be large enough to extracta FFOCT image. More precisely, applicants have measured the movements ofthe in vivo human eye and have shown that, when camera exposure time isset, for example, in a range of 1 ms to 10 ms (i.e. two consecutivecamera frames are acquired in 2-20 ms, respectively), the eye movementsinduced phase shift between the consecutive camera frames can take anyvalue from 0 to ±30 radians (or, equivalently, ±10π). More generally, invivo movements may induce phase changes between the consecutive directcamera images. These phase changes can be used to extract the FFOCTimage. According to this method FFOCT image can be extracted from the 2,4 or 5 direct images, but not restricted to this sequence, depending onthe scheme. Below, we will give example of FFOCT extraction method forthe 2 direct images, however this invention is not limited to 2 imagesscheme only, instead it is applicable to every FFOCT image retrievalscheme.

When the sample is moving along the Z direction, the phase of theinterference of the sample beam and the reference beam changes by arandom amount ψ. Different phase shifts may happen during the time thatcamera acquires an image. In the simplest case, it can be consideredthat each camera image has an average phase

ψ

. Then the recorded signal of the direct image on the camera is givenby:

${{I\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack {{\phi\left( {x,y} \right)} + \left\langle \psi \right\rangle} \right\rbrack}}} \right\}}},$

Where:

-   -   ϕ is the phase difference between the sample signal and the        reference signal;    -   ψ        is the random phase shift induced by the natural movements of        the in vivo sample. It is averaged over the acquisition time.    -   I₀ is the photon flux of the illumination;    -   R_(ref)(x, y)≈const is the reflectivity of the reference, which        is spatially uniform;    -   R_(sum)(x, y) is the reflectivity of the sample structures        within the coherence volume, which is the plane of interest;    -   R_(inc)(x, y) is the reflectivity of all the other structures        that are out of the coherence volume and other stray        reflections.

Then the two direct images are:

${I_{1}\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack {{\phi\left( {x,y} \right)} + \left\langle 0 \right\rangle} \right\rbrack}}} \right\}}$${I_{2}\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\cos\left\lbrack {{\phi\left( {x,y} \right)} + \left\langle \psi \right\rangle} \right\rbrack}}} \right\}}$

By subtracting two images and simplifying the formula we get:

${{I_{1}\left( {x,y} \right)} - {I_{2}\left( {x,y} \right)}} = {\frac{I_{0} \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}}}{2} \cdot \left\lbrack {\frac{\sin\;{\phi \cdot {\sin\left( \left\langle \psi \right\rangle \right)}}}{2} + {\cos\;{\phi \cdot {\sin^{2}\left( \frac{\left\langle \psi \right\rangle}{2} \right)}}}} \right\rbrack}$

From the formula, it can be seen that the FFOCT image can be obtainedfor every average phase difference

ψ

for the two consecutive or more distant camera frames, but maximum FFOCTsignal is achieved for

ψ

=π (considering that ϕ=0).

In FIG. 6, it is shown, that in order to get high FFOCT signal, not onlythe phase shift needs to be large enough, but also that it should happenduring the time, while camera acquires several frames (in FIG. 6 anexample of 2 frames is shown). By knowing the instantaneous speed of thein vivo movements of the sample, one can find the time interval that isneeded to achieve the average π phase shift. Stack of the direct imagesmay be recorded and the two frames, corresponding to the π phase shift,can be extracted from the stack and processed in order to get a highFFOCT signal. This method can be, for example, used for eye imaging:when the movement of the eye is such that the induced phase shiftbetween the two consecutive direct camera images is smaller than aradian (typically out of the large spikes liked to heart beat) then thephase-shift between these two images is sufficient to obtain an FFOCTimage. Additionally, in order to increase the number of useful (with πphase shift) direct images of the camera, one can adjust the cameraacquisition speed and the wavelength of the light source according tothe typical speeds of in vivo sample movements. Speed of the sample, atwhich the maximum FFOCT signal is achieved:

$\upsilon = \frac{\lambda \cdot \left\langle \psi \right\rangle}{2 \cdot \pi \cdot T}$

Where:

λ is the wavelength of the FFOCT light source.

T is the time that takes the camera to acquire two direct images

From the formula it follows that by initially knowing the typical speedof the sample in vivo movements ν, we can adjust the camera speed andthe wavelength of the light source to get the average π phase differencebetween the direct images (and therefore the best FFOCT signal) at thetypical speed of the sample. When the movement of the eye is such thatthe induced phase shift between two successive images is smaller than aradian (typically out of the large spikes liked to heart beat) two phaseimage of standard FFOCT is usable.

Previously, for simplicity purposes, it was considered that each cameraimage has an average phase

ψ

. It is possible to make a more comprehensive analysis by consideringthe phase at each moment of time ψ(t) and considering that cameraacquires the image by integrating the light during an exposure time (forexample, from time T₀ to time T₁).

${I\left( {i,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\int_{T_{0}}^{T_{1}}{\cos\left\lbrack {{\phi\left( {x,y} \right)} + {\psi(t)}} \right\rbrack}}}} \right\}}$

Then the two consecutive direct images are:

${I_{1}\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\int_{T_{0}}^{T_{1}}{{\cos\left\lbrack {{\phi\left( {x,y} \right)} + {\psi(t)}} \right\rbrack}dt}}}} \right\}}$${I_{2}\left( {x,y} \right)} = {\frac{I_{0}}{4} \cdot \left\{ {{R_{inc}\left( {x,y} \right)} + {R_{ref}\left( {x,y} \right)} + {2 \cdot \sqrt{{R_{sam}\left( {x,y} \right)} \cdot {R_{ref}\left( {x,y} \right)}} \cdot {\int_{T_{1}}^{T_{2}}{{\cos\left\lbrack {{\phi\left( {x,y} \right)} + {\psi(t)}} \right\rbrack}dt}}}} \right\}}$

By subtracting the two images and simplifying the formula we get:

I₁(x, y) − I₂(x, y) = const ⋅   [∫_(T₀)^(T₁)cos [ϕ(x, y) + ψ(t)]d t − ∫_(T₁)^(T₂)cos [ϕ(x, y) + ψ(t)]d t]

The applicants have measured the function ψ(t) for in vivo human eye andshown that high FFOCT signal can be reached for different cameraexposure times (for example, 1 ms-10 ms).

In the example of FIG. 6, in order to obtain FFOCT image we need toacquire at least two direct images, which have an average phase shiftbetween them equal to π (Y axis on the graph). If we know the averagemoving speed of the sample, we know the time that is needed for sampleto move by 7E (for example 3.4 ms, how it is shown on the graph). Thenwe can adjust the camera acquisition speed to acquire two images in 3.4ms time. As a result, phase-shifting is performed by the sample only.

FIG. 7 illustrates images of in-depth layers of the cornea acquiredusing consecutive two-dimensional interferometric signals which arephase shifted only by the natural eye movements, i.e. no move of thereference mirror is made. LED was emitting 850 nm wavelength light. Thelight spectrum was 30 nm wide, resulting in 7.8 μm thickness of theoptical slice. Camera was set to acquire 550 direct images per second.Each image was acquired by integrating the light during the exposuretime of 1.75 ms. During the exposure time the sample was moving andchanging the optical phase ψ(t) of the interferometric signals. Bysubtracting two consecutive direct images from the camera, we subtractthe two integrals over time-varying phase from the formula above andobtain the FFOCT images. The FFOCT signal depends on the function ψ(t)during the time from T₀ to T₂.

In FIG. 7, images 71 to 76 show respectively the reflection from theepithelium and tear film (71), the epithelium and sub-basal nerves (72),the anterior, middle and posterior stroma (73-75) and the endothelium(76) of in vivo human cornea.

FIG. 8 illustrates images 81, 82, 83, 84 of in-depth stromal layer ofthe cornea acquired using different camera exposure times (and,therefore, different camera frame rates of 550 frames/second, 300frames/second, 200 frames/second and 100 frames/second, respectively).The images are captured in the same conditions than in FIG. 7, i.e.using consecutive two-dimensional interferometric signals which arephase shifted only by the natural eye movements, i.e. no move of thereference mirror is made.

The applicants have shown that such embodiment enable to retrieve verygood quality images and considerably simplify the system without theneed of camera-piezo synchronization.

Although described by way of a number of detailed example embodiments,the systems and methods for in vivo, full-field interference microscopyimaging of a scattering three-dimensional sample according to thepresent description comprise various variants, modifications andimprovements that will be obvious to those skilled in the art, it beingunderstood that these various variants, modifications and improvementsfall within the scope of the invention such as defined by the followingclaims.

1. A method for in vivo, full-field interference microscopy imaging of ascattering three-dimensional sample, comprising: disposing the sample inan object arm of an interference device of a full-field OCT imagingsystem, wherein said interference device further comprises a referencearm with an optical lens and a first reflection surface; producing, ateach point of an imaging field, an interference between a reference waveobtained by reflection of incident light waves on an elementary surfaceof the first reflection surface corresponding to said point of theimaging field and an object wave obtained by backscattering of incidentlight waves by a voxel of a slice of the sample at a given depth, saidvoxel corresponding to said point of the imaging field, acquiring, usingan acquisition device of said full-field OCT imaging system, a temporalsuccession of two-dimensional interferometric signals resulting from theinterferences produced at each point of the imaging field; storing, foreach two-dimensional interferometric signal, a time of acquisition;providing, at each time of acquisition of the two-dimensionalinterferometric signals, cross-sectional images (X-Z) of both the sampleand said first reflection surface of said full-field OCT imaging systemusing an OCT imaging system; determining a plurality of en face images(X-Y) of a plurality of slices of the sample, each en face image beingdetermined from at least two two-dimensional interferometric signalshaving a given phase shift; determining from the cross-sectional imagesprovided by the OCT imaging system at the times of acquisition of eachof said two two-dimensional interferometric signals a depth (z) for eachen face image (X-Y) of said plurality of slices; determining a 3D imageof the sample from said plurality of en face images of said plurality ofslices of the sample and depths.
 2. The method according to claim 1,wherein said full-field OCT imaging system and said OCT imaging systembeing mounted on a moving platform, the method further comprises movingsaid platform at least along an optical axis (Z) of the object arm todetermine said plurality of en face images (X-Y).
 3. The methodaccording to claim 1, further comprising moving said platform at leastalong a direction (X, Y) perpendicular to said optical axis of theobject arm.
 4. The method according to claim 2, wherein said referencearm being mounted on a moving platform, the method further comprisesmoving said platform to compensate for defocus.
 5. The method accordingto claim 1, wherein said object arm being mounted on a moving platform,the method further comprises moving said platform along an optical axis(Z) of the object arm to determine said plurality of en face images(X-Y).
 6. The method according to claim 5, wherein said reference armbeing mounted on a moving platform, the method further comprises movingsaid platform to compensate for defocus.
 7. The method according toclaim 1, wherein said reference arm being mounted on a moving platform,the method further comprises moving said platform along an optical axis(X) of the reference arm to compensate for defocus, to determine saidplurality of en face images (X-Y).
 8. The method according to claim 1,further comprising position shifting said first reflection surface ofthe reference arm of the full-field OCT imaging system to provide saidphase shift between said at least two two-dimensional interferometricsignals.
 9. The method according to claim 1, further comprisingselecting in said temporal succession of two-dimensional interferometricsignals acquired by the acquisition device, said at least twotwo-dimensional interferometric signals having said phase shift, whereinthe phase shift results from in vivo movements of the sample.
 10. Asystem for in vivo, full-field interference microscopy imaging of ascattering three-dimensional sample comprising: a full-field OCT imagingsystem for providing en face images of the sample, wherein saidfull-field OCT system comprises: an interference device comprising anobject arm intended to receive the sample and a reference arm comprisingan optical lens and a first reflection surface, wherein said object armand said reference arm are separated by a beam splitter and wherein theinterference device is adapted to produce, when the sample is disposedon the object arm of the interference device, at each point of animaging field, an interference between a reference wave obtained byreflection of incident light waves on an elementary surface of the firstreflection surface corresponding to said point of the imaging field andan object wave obtained by backscattering of incident light waves by avoxel of a slice of the sample at a given depth, said voxelcorresponding to said point of the imaging field, an acquisition deviceconfigured to acquire a temporal succession of two-dimensionalinterferometric signals resulting from the interferences produced ateach point of the imaging field, an OCT imaging system for providing atthe same times of acquisition of said two-dimensional interferometricsignals, cross-sectional images of both the sample and said firstreflection surface of said full-field OCT imaging system; a processingunit configured to: determine a plurality of en face images (X-Y) of aplurality of slices of the sample, each en face image being determinedfrom at least two two-dimensional interferometric signals having a givenphase shift; determine from the cross-sectional images provided by theOCT imaging system at the times of acquisition of each of said twotwo-dimensional interferometric signals a depth (z) for each en faceimage (X-Y) of said plurality of slices; determine a 3D image of thesample from said plurality of en face images of said plurality of slicesof the sample and depths.
 11. The system according to claim 10, whereinsaid first reflection surface of the reference arm of the full-field OCTimaging system is position shifted to provide said optical pathdifference between said at least two-dimensional interferometricsignals.
 12. The system according to claim 10, wherein the processingunit is further configured to: select in said temporal succession oftwo-dimensional interferometric signals acquired by the acquisitiondevice, said at least two-dimensional interferometric signals havingsaid given optical path difference, wherein the optical path differenceresults from in vivo movements of the sample.
 13. The system accordingto claim 10, wherein said object arm of the full-field OCT imagingsystem further comprises an optical lens.
 14. The system according toany of claim 10, wherein said reference arm and/or object arm of thefull-field OCT imaging system can be moved with respect to said beamsplitter of the interference device of said full-field OCT imagingsystem.
 15. The system according to claim 10, further comprising amoving platform, wherein said full-field OCT imaging system and said OCTimaging system are mounted on said moving platform.
 16. The systemaccording to claim 10, wherein the OCT imaging system is a spectraldomain OCT imaging system, a time-domain OCT imaging system, or aswept-source OCT imaging system.